An MRI system is used to perform imaging on medical patients to ascertain details regarding internal structures. Typically, a portion of a patient is positioned within a large magnet. In addition to the main magnet, gradient magnets are further adapted within the MRI system to generate gradient fields. Then radio frequency (RF) energy is applied to a coil within the magnet to cause precession of protons within the patient based on the magnetic field conditions. When the RF energy is turned off, movement of the protons releases energy, which generates a signal that can be received and processed.
Main magnetic field (B0) inhomogeneity is the fractional deviation of the local magnetic field from its average value over a specified diameter spherical volume (DSV). Poor main (B0) magnetic field homogeneity (MFH) of an MRI system leads to artifacts and signal losses in MR images. Technical difficulties of creating a perfectly uniform field can cause poor MFH. Also, susceptibility differences within objects being imaged are a major source of MFH. Clinical application of magnetic resonance spectroscopy (MRS) data may be rejected if voxels are located near the sinuses or temporal bones and where poor MFH is demonstrated. Poor MFH also may lead to geometrical distortions of MR images or broaden the spectral linewidths in MRS. These distortions can be either spatial distortions, intensity nonuniformities or both. MRI pulse sequences that are sensitive to magnetic homogeneity include echo-planar imaging (EPI), true fast imaging by steady precession (trueFISP) and fast field echo (FFE) imaging methods. In cardiac MRI, gross motion, blood flow, and chemical shift artifacts from epicardial fat must all be managed, and thus a homogeneous static magnetic field is imperative.
It is often difficult to measure MFH for clinically deployed MRI systems in a manner that is time-efficient. For example, as part of an installation procedure, MRI service personnel may spend many man-hours evaluating MFH by placing a small sample in the magnet and determining its resonant frequency at hundreds of different locations. However, this process is time consuming and is only performed on installation.
The American College of Radiology (ACR) in its MRI Accreditation Program's Quality Control (QC) Manual describes two methods for measuring MFH: the spectral linewidth method (FWHM) and the phase difference (Δφ) method.
For the linewidth method, the MFH is measured from the full width at half maximum (FWHM) of water peak of a phantom. The FWHM is converted from frequency in Hertz (Hz) to parts per million (ppm) using the following equation (for a proton spectrum):FWHM(ppm)=FWHM(Hz)/42.576B0(T)  [EQ.1]where B0 is the main magnetic field strength. However, this method is only suitable for measuring MFH at a single DSV.
For the phase-difference (Δφ) method, homogeneity maps are generated using a FFE sequence to acquire phase Fourier transformed images using two time-to-echo values, TE1 and TE2. Unwrapping and subtraction of the two-phase images is accomplished by a special image processing routine that is not generally accessible to a system operator. According to the ACR's MRI QC Manual, the greatest difference of ΔB0 divided by B0 will give the MFH in ppm. Let ΔB0 (mT)=B0(0)−B0(r), where B0(0) is measured at the isocenter and B0(r) is measured at a pixel located a distance, r, from the isocenter. Then ΔB0 (mT) can be calculated by using the phase difference Δφ (in radians) divided by the gyromagnetic ratio (γ=267,513 radians/mT for protons) and multiplied by the difference of the inverse of TE1 and TE2 in units of seconds as follows:
                              Δ          ⁢                                          ⁢                      B            0                          =                                            ∂              φ                        r                    ⁢                      (                                          1                                  TE                  ⁢                                                                          ⁢                  1                                            -                              1                                  TE                  ⁢                                                                          ⁢                  2                                                      )                                              [                  EQ          .                                          ⁢          2                ]            The Δφ method has found particular favor in research facilities that routinely perform functional MRI studies using blood-oxygen level dependent contrast. However, this method is often unavailable on a given MRI system.
Both of these methods are time-consuming, require significant operator expertise, and cannot be universally used in available MRI systems. Furthermore, these known methods can provide only limited information regarding MFH.
Accordingly, a need exists for an improved method of performing MFH that is suitable across MRI systems.